Ultrasonic perfusion measurement using contrast agents

ABSTRACT

Apparatus and methods are disclosed for the detection and imaging of ultrasonic contrast agents. Ultrasonic apparatus is provided for coherent imaging of ultrasonic contrast agents, and for detecting harmonic contrast agents. The inventive apparatus includes a dual display for simultaneously viewing a real time image which displays the location of the contrast agent and a triggered contrast image. Methods of contrast agent detection and imaging include the measurement of perfusion rate characteristics, multizone contrast imaging, multifrequency contrast imaging, tissue perfusion display, and high PRF contrast image artifact elimination.

[0001] This invention relates to ultrasonic diagnosis and imaging of thebody with ultrasonic contrast agents and, in particular, to new methodsand apparatus for ultrasonically detecting and imaging with contrastagents.

[0002] Ultrasonic diagnostic imaging systems are capable of imaging andmeasuring the physiology within the body in a completely noninvasivemanner. Ultrasonic waves are transmitted into the body from the surfaceof the skin and are reflected from tissue and cells within the body. Thereflected echoes are received by an ultrasonic transducer and processedto produce an image or measurement of blood flow. Diagnosis is therebypossible with no intervention into the body of the patient.

[0003] Materials known as ultrasonic contrast agents can be introducedinto the body to enhance ultrasonic diagnosis. Contrast agents aresubstances which strongly interact with ultrasonic waves, returningechoes which may be clearly distinguished from those returned by bloodand tissue. One class of substances which has been found to beespecially useful as an ultrasonic contrast agent is gases, in the formof tiny bubbles called microbubbles. Microbubbles present a significantacoustic impedance mismatch in the body, and nonlinear behavior incertain acoustic fields which is readily detectable through specialultrasonic processing. Gases that have been stabilized in solutions inthe form of tiny microbubbles are infused into the body and survivepassage through the pulmonary system and circulate throughout thevascular system. Microbubble contrast agents are useful for imaging thebody's vascular system, for instance, as the contrast agent can beinjected into the bloodstream and will pass through the veins andarteries of the body with the blood supply until filtered from the bloodstream in the lungs, kidneys and liver.

[0004] One type of microbubble contrast agent currently underinvestigation comprises coated microbubbles. The microbubbles of thecontrast agent are covered with a thin biodegradable coating or shell.The microbubbles have diameters between 0.1 μm and 4.0 μm and a specificdensity about {fraction (1/10)} of the density of water. The coatedmicrobubbles are suspended in an aqueous solution for infusion into theblood stream.

[0005] Coated microbubbles have the advantage of being stable in thebody for a significant period of time, as the shells serve to protectthe gases of the microbubbles from diffusion into the bloodstream. Thesize of the microbubbles is chosen to enable the microbubbles to passthrough capillary beds in the body.

[0006] At moderately high sound pressure amplitudes the acousticpressure waves can cause the shells of coated microbubbles to rupture,freeing the bubbles to behave as noncoated microbubbles until theydiffuse into the bloodstream. In their noncoated form acoustic energycan induce nonlinear motion of the microbubbles, itself a detectableultrasonic phenomenon. This acoustically induced destruction andcollapse of the microbubbles produces a high amplitude response and acharacteristically bright pattern in the color Doppler mode. Hence colorDoppler is an advantageous modality for detecting the collapse ofcontrast agent microbubbles.

[0007] U.S. Pat. No. 5,456,257, assigned to the same assignee as thepresent invention, describes a technique for detecting microbubblesthrough phase insensitive detection of microbubble destruction anddifferentiation of the detected signals on a spatial basis. Phaseinsensitive contrast agent detection advantageously reduces artifactsfrom moving tissue, and also performs well when imaging contrast agentperfused tissue, where the contrast agent is finely distributed andmoving slowly through the fine capillary structure of tissue. It isdesirable to be able to perform contrast agent imaging with equal effectin large, rapidly moving blood pools such as the chambers of the heart.It is also desirable to specifically tailor the operation of theultrasound machine to harmonic characteristics when performing harmoniccontrast imaging.

[0008] In accordance with the principles of present invention, new andimproved apparatus and methods for the detection and imaging ofultrasonic contrast agents are provided. Ultrasonic apparatus isprovided for coherent imaging of ultrasonic contrast agents, which isadvantageous in blood pool contrast imaging. In a second embodiment, theapparatus is specially tailored to be programmed with responsecharacteristics suitable for harmonic contrast agents. The inventiveapparatus also includes a display for simultaneously viewing a real timeimage which displays anatomical structures for localization of thecontrast agent and a triggered contrast image displaying contrastenhanced images. Methods of employing the inventive apparatus withcontrast agents include the measurement of perfusion ratecharacteristics, multizone contrast imaging, a technique for discerninglarger vessels in a bed of fine capillary structures, multifrequencycontrast imaging, the display of contrast enhanced tissue, and atechnique for the elimination of artifacts occurring during high PRFcontrast image acquisition.

[0009] In the drawings:

[0010]FIG. 1 illustrates in block diagram form an ultrasonic diagnosticsystem of U.S. Pat. No. 5,456,257 which is capable of performing phaseinsensitive contrast agent detection;

[0011]FIG. 2 illustrates in block diagram form an ultrasonic diagnosticsystem of the present invention which is capable of performing coherentcontrast agent detection;

[0012]FIG. 3 illustrates an ultrasonic image display for contrast agentimaging;

[0013]FIG. 4 illustrates in block diagram form a second embodiment ofthe present invention which provides performance advantages for harmoniccontrast agent detection;

[0014]FIGS. 5 and 6 illustrate passband characteristics used to explainthe performance of the embodiment of FIG. 4;

[0015]FIG. 7 illustrates the principle of time separated pulsing whenimaging contrast agents;

[0016]FIG. 8 illustrates an FIR filter structure suitable for use in theembodiment of FIG. 4;

[0017]FIGS. 9a-9 d illustrate the effects of stenoses on contrast agentperfusion;

[0018]FIGS. 10a and 10 b illustrate perfusion curves for good and poorperfusion rates;

[0019]FIGS. 11a and 11 b illustrate repetitive perfusion curves for goodand poor perfusion rates;

[0020]FIG. 12 illustrates a triggering technique for estimating theperfusion curve of FIG. 13;

[0021]FIGS. 14a-14 c illustrate a multizone contrast agent scanningtechnique;

[0022]FIGS. 15a and 15 b illustrate display mapping characteristics forcontrast agent imaging;

[0023]FIG. 16 illustrates the heart in cross section; and

[0024]FIGS. 17a-17 c illustrate the removal of artifacts occurringduring high PRF contrast imaging.

[0025] Referring to FIG. 1, an ultrasonic diagnostic system described inU.S. Pat. No. 5,456,257 is illustrated in block diagram form. Thisultrasound system is capable of performing phase insensitive contrastagent detection as described in that patent. In the illustrated system,coherent echo signals produced by a beamformer 16 are quadraturedemodulated by an I,Q demodulator 18 to produce quadrature I and Qsignal components. The demodulated signal components are amplitudedetected by an envelope detector 20. The detected signals are filteredby a filter 22 to remove noise and other extraneous signal components.Spatially aligned, temporally separated detected echo signals aredifferentiated by a pulse to pulse differentiation subsystem 24, and thedifferential signals are used to form contrast agent enhanced images.

[0026] Performing pulse to pulse differentiation of envelope detectedecho signals provides advantages in certain procedures. When contrastagents are being used in a mode where microbubbles in the fine capillarystructures of tissue are destroyed by ultrasonic waves, differentiationof the echo envelope is particularly useful. In this mode of operation afirst ultrasonic pulse destroys the microbubbles in the tissue and thesedestruction events are received and envelope detected. A second pulse istransmitted to the same locations, and the returning echoes, ideally,show an absence of microbubbles at the locations where the microbubbleswere destroyed. The second set of echoes is subtracted from the firstset on a spatial basis, yielding difference signals of substantialmagnitude at the locations where the microbubbles were destroyed, whichare then displayed at corresponding pixel locations on a display. In arealistic setting the second set of echoes may not actually reveal voidswhere microbubbles were destroyed, due to motional effects, diffusionrates, and other bubble activity. However the difference in bubbleactivity from one pulse to the next will provide a highly detectableresponse when differentiated on a pulse to pulse, spatial basis.

[0027] This dramatic difference in scattering characteristics ofmicrobubbles from one pulse to another may be due to a host of factors:the bursting of microbubble coatings; oscillation and nonlinearmicrobubble motion; diffusion of a microbubble during the interpulseinterval; or bubble repositioning, for instance. When microbubbledestruction is referred to herein it encompasses the effects ofphenomena such as these.

[0028] A first embodiment of an ultrasonic diagnostic system constructedin accordance with the principles of the present invention is shown inFIG. 2. This embodiment provides coherent detection of ultrasoniccontrast agents. An ultrasonic probe 10 includes an array 12 ofultrasonic transducers which transmit and receive ultrasonic energy.During transmission, an ultrasonic beamformer 16 controls the timing ofactuation of the separate elements of the array 12 by activating thetransducer pulsers of a transmitter/receiver 14 at appropriate times topulse the transducer elements so that a steered and focused ultrasonicbeam is produced. During reception ultrasonic echoes received by thetransducer elements are received by transmitter/receiver 14 and coupledto separate channels of the beamformer 16, where the signals areappropriately delayed then combined to form a sequence of coherent echosignals over the depth of reception in the body of the patient.

[0029] The coherent echo signals are quadrature demodulated by an I,Qdemodulator 18 which produces quadrature I and Q signal components. Thedemodulated signal components are coupled to a B mode processor 37 whichfilters, detects, and maps greyscale echo signals in the usual manner.The greyscale echo signals of the scanlines of an image are coupled to ascan converter 40 for display of a B mode image.

[0030] In accordance with the principles of the present invention the Iand Q signal components are alternatively (or in addition) coupled to apulse to pulse differentiation circuit 24 which differentiates echoesreceived from the same sample volume (location) in the body on atemporal basis. The results of this differentiation are coupled to anamplitude detector 20 and the differential response signals are coupledto an event discriminator 27. The event discriminator discriminatesevents of microbubble destruction at the sample volume location from thedifferentiated echo information. One convenient way to perform thisdiscrimination is by comparison of the detected signals to a thresholdfrom threshold generator 26, passing signals above a threshold andrejecting signals below the threshold. The discriminator will detectmicrobubble destruction events and reject low level noise.

[0031] Detected events are coupled to the scan converter 40 forproduction of a spatial image of the microbubble destruction events inthe desired image format. The destruction event image may be shownseparately, or may be combined with the B mode image to show thecontrast agent in relation to surrounding tissue structure. The imagesare coupled to a video processor 42 which produces video signals fordisplay on an image display 50.

[0032] This coherent contrast agent detection technique is highlysensitive to small variations in microbubble activity in the image area,and performs well when imaging the blood pool of a heart chamber, forinstance. In a large blood pool with a large population of movingmicrobubbles, the probability of differential microbubble activity fromone pulse to another is extremely high, explaining the high sensitivityof this technique for heart chamber imaging. In comparison with theincoherent contrast agent detection technique, coherent microbubbledetection is more sensitive to tissue motion and more sensitive toindividual microbubble events in high contrast agent concentrations. Itis also possible to process received echoes both coherently andincoherently, to form an image which contains information from bothprocesses.

[0033] Contrast agent detection in accordance with the present inventionprovides excellent tissue clutter rejection. Microbubble echo signalsare generally not received alone, but are usually accompanied by echosignals of much greater amplitude which are returned from neighboringtissue and structures. These tissue echoes can be several orders ofmagnitude greater than any of the microbubble echo signals, effectivelymasking them. The pulse to pulse differentiation processing can rejectthe tissue signals by effectively canceling them, revealing the contrastagent echoes which can then be more readily discriminated. Thiscancellation is enhanced by high PRF pulses, which further diminishesmotional artifacts from tissue.

[0034] It has been found that the microbubbles of a contrast agentexhibit greater sensitivity to ultrasonic pulses of certaincharacteristics and lesser sensitivity to pulses of othercharacteristics. In general, the higher the amplitude, the lower thefrequency, and (to a lesser extent) the longer the burst length, themore sensitive the microbubbles are to destruction. Thus, the time ofoccurrence of microbubble destruction can be modulated and controlled.Microbubbles can be imaged in the bloodstream by scanning at a highfrequency with low amplitude, (and to a lesser extent) short burstlength pulses. When it is desired to stimulate microbubble destruction,higher power pulses of lower frequency and longer burst length aretransmitted into the bloodstream. The ultrasound system of the presentinvention is provided with control presets for these two pulsetransmission characteristics, enabling the clinician to switch fromnondestructive imaging pulses to microbubble destruction pulses when theclinician so desires. The preferred display of anatomical structures andmicrobubble activity employs programmed switching between thedestructive and nondestructive pulse modes for contrast agent andanatomical structure imaging.

[0035] A second embodiment of an ultrasonic diagnostic systemconstructed in accordance with the principles of the present inventionfor use with harmonic contrast agents is shown in FIG. 4. In this secondembodiment the array transducer 112 of the probe 110 transmitsultrasonic energy and receives echoes returned in response to thistransmission. The response characteristic of the transducer can exhibittwo passbands, one around the central transmit frequency and anotherabout the center of the received passband. For imaging harmonic contrastagents, a broadband transducer having a passband encompassing both thetransmit and receive passbands is preferred. The transducer may bemanufactured and tuned to exhibit a response characteristic as shown inFIG. 5, in which the lower hump 60 of the response characteristic iscentered about the center transmit frequency f_(t), and the upper hump62 is centered about the center frequency f_(r) of the responsepassband. The transducer response characteristic of FIG. 6 is preferred,however, as the single dominant characteristic 64 allows the probe to besuitable for both harmonic contrast imaging and imaging without harmoniccontrast agents. The characteristic 64 encompasses the central transmitfrequency f_(t), and also the harmonic receive passband bounded betweenfrequencies f_(L) and f_(c), and centered about frequency f_(r). Atypical harmonic contrast agent can have a response such thattransmission about a central transmit frequency of 1.7 MHz will resultin harmonic returning echo signals about a frequency of 3.4 MHz. Aresponse characteristic 64 of approximately 2 MHz would be suitable forthese harmonic frequencies.

[0036] In FIG. 4 a central controller 120 provides a control signalf_(tr) to a transmit frequency control circuit 121 to control the centerfrequency and time of transmission of the transmitted ultrasonic energy.The transmit frequency control circuit pulses the elements of thetransducer array 112 by means of a transmit/receive switch 114. Apreferred method of pulsing the transducer array is in bursts which scanwith sufficient pulses to form an image, followed by intervals of nopulse transmission. Such bursts and intervals are shown in FIG. 7, whichshows a burst interval nPRF and a frame interval t_(Fr), the frameinterval including the burst interval and an interval of no pulsetransmission. The latter interval allows time for new contrast agentcoursing through the body to infuse the vessels and tissue of the imageplane between frame bursts. The frame intervals can be on the order ofone second, and can be gated to the heart rate or asynchronous withrespect to the heart rate. During each nPRF burst interval, echoes fromthe same spatial locations can be gathered for Doppler processing.Preferably a high PRF rate such as 6 KHz is used. Imaging procedures ofthis type are the subject of U.S. patent [application Ser. Nos.08/439,619 and 08/540,463].

[0037] Medical diagnostic ultrasonic scanning is limited by regulatoryrequirements in the peak pressure amplitude of a transmitted pulse andthe integral of the energy transmitted. The preferred scanning ofcontrast agents in accordance with the embodiment of FIG. 4 utilizesrelatively high peak pulse power, with the time integral of transmittedenergy lessened by the intervals during which no pulses are transmitted.The ultrasound system is set to operate with a relatively highmechanical index and an SPTA moderated by the gated or interval bursts.

[0038] Echoes received by the transducer array 112 are coupled throughthe T/R switch 114 and digitized by analog to digital converters 115.The sampling frequency f_(s) of the A/D converters 115 is controlled bythe central controller. The desired sampling rate dictated by samplingtheory is at least twice the highest frequency f_(c) of the receivedpassband and, for the preceding exemplary frequencies, might be on theorder of at least 8 MHz. Sampling rates higher than the minimumrequirement are also desirable.

[0039] The echo signal samples from the individual transducer elementsare delayed and summed by a beamformer 116 to form coherent echosignals. The digital coherent echo signals are then filtered by adigital filter 118. In this embodiment, the transmit frequency f_(t) isnot tied to the receiver, and hence the receiver is free to receive aband of frequencies which is separate from the transmitted band. Thedigital filter 118 bandpass filters the signals in the passband boundedby frequencies f_(L) and f_(c) in FIG. 6, and can also shift thefrequency band to a lower or baseband frequency range. The digitalfilter could be a filter with a 1 MHz passband and a center frequency of3.4 MHz in the above example. A preferred digital filter is a series ofmultipliers 70-73 and accumulators 80-83 as shown in FIG. 8. Thisarrangement is controlled by the central controller 120, which providesmultiplier weights and decimation control which control thecharacteristics of the digital filter. Preferably the arrangement iscontrolled to operate as a finite impulse response (FIR) filter, andperforms both filtering and decimation. For example, only the firststage output 1 could be controlled to operate as a four tap FIR filterwith a 4:1 decimation rate. Temporally discrete echo samples S areapplied to the multiplier 70 of the first stage. As the samples S areapplied, they are multiplied by weights provided by the centralcontroller 120. Each of these products is stored in the accumulator 80until four such products have been accumulated (added). An output signalis then produced at the first stage output 1. The output signal has beenfiltered by a four tap FIR filter since the accumulated total comprisesfour weighted samples. Since the time of four samples is required toaccumulate the output signal, a 4:1 decimation rate is achieved. Oneoutput signal is produced for every four input samples. The accumulatoris cleared and the process repeats. It is seen that the higher thedecimation rate (the longer the interval between output signals), thegreater can be the effective tap number of the filter.

[0040] Alternatively, temporally separate samples are delayed by delayelements T and applied to the four multipliers 70-73, multiplied, andaccumulated in the accumulators 80-83. After each accumulator hasaccumulated two products, the four output signals are combined as asingle output signal. This means that the filter is operating as aneight tap filter with a 2:1 decimation rate. With no decimation, thearrangement can be operated as a four tap FIR filter. The filter canalso be operated by applying echo signals to all multiplierssimultaneously and selectively time sequencing the weightingcoefficients. A whole range of filter characteristics are possiblethrough programming of the weighting and decimation rates of the filter,under control of the central controller.

[0041] Returning to FIG. 4, filtered echo signals from tissue, generallyfiltered by a passband centered about or demodulated from the transmitfrequency, are coupled to a B mode processor 37 for conventional B modeprocessing. Filtered echo signals of the contrast agent passband arecoupled to a contrast signal detector 128 which eliminates stationarytissue signals by pulse to pulse subtraction of temporally discreteechoes from a given spatial location, amplitude or envelope detects theresulting difference signals, and discriminates for motion signalcomponents on an amplitude basis. Simple two pulse subtraction of theform P1-P2 may be employed where P1 represents the echoes receivedfollowing one pulse and P2 represents the echoes received followinganother pulse. Three pulse subtraction of the form |P₁−P₂|+|P₂−P₃| maybe employed to accumulate more signals from successive bubbledestruction pulses.

[0042] The filtered echo signals from the digital filter 118 are alsocoupled to a Doppler processor 130 for conventional Doppler processingto produce velocity and power Doppler signals. The outputs of theseprocessors are coupled to a 3D image rendering processor 132 for therendering of three dimensional images, which are stored in a 3D imagememory 134. Three dimensional rendering may be performed as described inU.S. patent [application Ser. No. 08/638,710], and in U.S. Pat. Nos.5,474,073 and 5,485,842, the latter two patents illustrating threedimensional power Doppler ultrasonic imaging techniques. The signalsfrom the contrast signal detector 128, the processors 37 and 130, andthe three dimensional image signals are coupled to a video processor 140where they may be selected for display on an image display 50 asdictated by user selection. The video processor preferably includespersistence processing, whereby momentary intensity peaks of detectedcontrast agents can be sustained in the image. One technique forproviding persistence is through frame averaging, whereby new imageframes are combined with previous frame information on a spatial basis.The combination can be done by weighting the contributions of the oldand new frame information and the frame information can be combined in arecursive manner; that is, old frame information is fed back forcombining with new frame information. A preferred persistence techniqueis the fast attack, slow decay technique described in U.S. Pat. No.5,215,094, which can be applied to both Doppler and contrast agentimages.

[0043] Several imaging formats have been found to be preferred forcontrast imaging. Power motion imaging as described in U.S. patent[application Ser. No. 08/655,391] in which the intensity of signalsresulting from moving tissue is displayed, has been found to be highlydiagnostic for structures such as the walls of the heart when perfusedwith contrast agents. Power Doppler imaging has been found to yieldexcellent results for bloodflow. Three dimensional power Doppler imagingof vessels infused with contrast agent provide excellent visualizationof the continuity of bloodflow and stenoses. The combination of B modeor power motion structural information with power Doppler signals inaccordance with the semi-transparent rendering techniques of theaforementioned patent [application Ser. No. 08/638,710] provides superbrenderings of both flow and surrounding structure.

[0044] A preferred display format for contrast agent imaging is depictedby the screen display of FIG. 3. In this display the signals produced bythe B mode processor 37 are used to display a real time image display160 of structure in the body such as a blood vessel 170. This real timeimage is used by the clinician to ascertain and locate the area of thebody to be imaged. Preferably the B mode image is created from echoesreturning from nondestructive ultrasonic imaging pulses. As discussedabove, pulses of low amplitude, high frequency, and short burst durationwill generally not destroy the microbubbles. However, echoes from pulsesdestructive of microbubbles are used by the contrast signal detector 128to produce contrast agent images 160′ on the same or an adjacentmonitor. Preferably the contrast agent images 160′ are triggered to beacquired at a predetermined phase of the heart cycle, using a heart gatetriggering from the phases of the heartbeat waveform. When the heartbeatis at the desired phase of its cycle, a burst of relatively highamplitude, low frequency, long burst duration pulses are transmitted todestroy the microbubbles in the image plane and detect and display thoseevents. A B mode image acquired at or near the same heartbeat phase isdisplayed, with the vessel or organ 170′ filled in with the imagedmicrobubble destruction events. Thus, the display screen of FIG. 3 willshow a B mode image 160 in real time, and a contrast agent image 160′which is updated each heart cycle.

[0045] While the foregoing image presentation is especially useful incardiology where the beating heart is constantly in motion, a variationof this presentation is especially useful in radiology where tissuestructure is more stationary. In the variation, a real time B mode image160′ of anatomical structure is shown, with fluid flow 170′ filled inwith color Doppler. This real time color flow Doppler image is thenperiodically filled in with detected contrast agent, sharplyilluminating the bloodflow. The colorflow Doppler display and thecontrast agent display, both of which are filling in the same areas ofthe anatomical display, may be shown in the same, similar, orcontrasting colors and intensities. The periodicity of the overlaidcontrast agent display may be synchronized to the heart cycle with anEKG trigger as described above, or the periodicity may be chosen by theuser and asynchronous to the heart cycle.

[0046] A contrast agent procedure which is advantageously performed inaccordance with the present invention is the measurement of the rate ofperfusion of an organ or area of the body. FIG. 9a illustrates thetravel of an intravenous injection of contrast agent to a capillary bed200. The agent travels in the bloodstream as it moves from the injectionsite 208 and traverses the right ventricle 202, the lungs 204, and theleft ventricle 206 before reaching an artery 209. The contrast agentthen begins to infuse the tissue of the capillary bed 200 as blood flowsfrom the artery 200 through the arterioles 210 and into the capillariesof the tissue.

[0047] The perfusion rate into the capillary bed can be used to evaluatethe viability of bloodflow in that region of the body or to identify thelocation of a stenosis. Ultrasonic pulses are transmitted to destroymicrobubbles in a region 212 across the capillary bed 200, as shown inFIG. 9b. If a stenosis 214 is impeding the flow of blood in the artery209 and hence to the entire capillary bed 200, the rate of reperfusionof microbubbles will be slow across the entire region 212. But if thestenosis 216 is in an artery which feeds only part of the capillary bed200, the rate of perfusion will be slow in only the portion 218 of theregion which is fed through the stenotic artery. This difference in therate of reperfusion is illustrated graphically by the curves of FIGS.10a and 10 b. Each of these curves shows the same blood volume and hencethe same initial microbubble concentration 220 before the microbubblesare destroyed in the capillary bed. At time td ultrasonic pulses destroythe microbubbles as indicated by the vertical spike in each curve. Whenblood is flowing freely into the capillary bed, a rapid rate ofreperfusion of microbubbles occurs as indicated by curve 222 in FIG.10a. The curve 222 rapidly rises back to the stable microbubbleconcentration level 220. But when the bloodflow is impeded, the rise ofthe curve 224 is much more gradual, as indicated in FIG. 10b. Thereperfusion curve can be repeated continually as shown by FIGS. 11a and11 b. FIG. 11a shows a repetitive sequence of reperfusion curves 222,each returning to the full perfusion level 220 in a period of timet_(p). In FIG. 11b, each curve 224 of the same duration t_(p) is shortof the full perfusion level 220 by an amount indicated by arrows B-B.

[0048] The reperfusion curve may be reproduced as indicated in FIG. 13.Ultrasonic pulses are transmitted at time td to destroy the microbubblesin the capillary bed. A short time later pulses are transmitted again,the echoes received and imaged to this time measure the degree ofmicrobubble reinfusion, either by destroying reinfused microbubbles andrecording the destruction events, or by counting or integrating pixelsin the area which show reinfused microbubbles. The measure of the numberof microbubbles reinfused to the region is plotted as a point X of thecurve 224. Nondestructive pulses can be repetitively transmitted andechoes received to plot a sequences of X points on the curve as shown inFIG. 13.

[0049] Another way to measure the X points on the reinfusion curvethrough readily detectable microbubble destruction events is to utilizea cyclic measure similar to the repetitive pattern of FIG. 11b. Thecyclic measure is useful where the flow in the region is stronglypulsatile due to the heartbeat cycle. FIG. 12 shows a heart cyclewaveform 230, indicating the pulsatile action of bloodflow. At the peaksof the waveform 230, new blood is pumped into regions of the body duringthe systolic phase of the heart cycle. Advantage is taken of thisreinfusing action by repetitively measuring the degree of contrast agentreinfusion at a constant point in the heart cycle, but followingcontinually differing phases of microbubble destruction. In FIG. 12 theX points of reinfusion measurement all occur at the same phase of theheart cycle. The X points are preceded by changing times at which themicrobubbles are destroyed, as indicated by arrows 232, 234, and 236,which successively precess to earlier times in the heart cycle. Thismeans that each X_(n) point of FIG. 12 will be a later X_(n) point onthe curve 224 of FIG. 13. Since the purpose of ultrasonic transmissionat the times of arrows 232, 234, and 236 is to destroy the microbubbles,it is not necessary to receive and analyze the returning echoes at thesetimes. Echo reception and analysis is done at the times of the Xs, andthe Xs shown in FIG. 12 can be plotted as the successive Xs in FIG. 13due to the precession of the destruction time phases indicated by thearrows.

[0050] For cardiac imaging it may be desirable to trigger the X_(n)times in synchronization with the diastolic phase of the heart cyclewhen the coronary arteries are reinfused with blood. Triggered or gatedacquisition is especially significant in cardiac imaging to reducetissue motion artifacts stemming from the beating movement of the heart.

[0051] This technique of measuring perfusion by microbubble destructioncan also be used to image the flow in major vessels of a capillary bed.In FIG. 9d, for instance, it is seen that the major vessels 240 reinfuseearlier than the fine capillaries in a microbubble depleted region 212.The major vessels 240 can be revealed by detecting microbubbles in theregion 212 shortly after pulses have destroyed all of the microbubblesin the region, at which time only the major vessels 240 have beensignificantly reinfused with contrast agent.

[0052] It has been found that it is at times not possible to destroy allmicrobubbles in the image plane due to several factors. Sincemicrobubbles are destroyed by high energy, focused ultrasound beams tendto destroy more microbubbles near the beam focal point than at otherlocations. Also, when a dense concentration of microbubbles is to bedestroyed, a great deal of the ultrasonic pulse energy is attenuated bythe near field microbubbles, leaving insufficient energy to destroy farfield microbubbles. A technique for overcoming these effects is shown inFIGS. 14a-14 c. In these drawings, the horizontal axis represents depthinto the body, with the skin line SL indicated at the left side of eachdrawing. A typical ultrasonic image may show the skin line at the top ofthe image and the deepest penetration into the body at the bottom of theimage. To bring a maximum level of energy to bear on the microbubbles inthe image plane, focused pulses are transmitted to focus the ultrasonicenergy on the microbubbles which are to be destroyed. When imaging is tobe done to a significant depth in the body, the pulses will not befocused over the full image depth, but will come into focus around aparticular focal point and then diverge at greater depths. This isindicated in FIG. 14a, where a transmitted pulse is focused at a focalpoint F₁ which is in a focal zone Z₁.

[0053] Above this first focal zone Z₁ is a line 270, which representscomplete microbubble destruction over this near field part of the focalzone Z₁ and about the focal point F₁. Beyond the focal point the degreeof microbubble destruction decreases, as indicated by the declining line272. These lines are shown as straight lines for ease of explanation; itwill be understood that the effect will usually be continually changingand that actual effects may follow a curved relationship.

[0054]FIG. 14a represents the transmission of a first pulse along agiven beam direction, a result of which is that near field microbubblesare destroyed as indicated by lines 270 and 272. Following thismicrobubble destruction, a second pulse is transmitted to gather echoesfrom along the microbubble depleted beam direction. The echoes from thetwo pulses may be differentiated and displayed using the ultrasonicapparatus of FIGS. 1, 2, or 4.

[0055] The next pulse transmission for microbubble destruction isfocused at a second focal point F₂ in a second focal zone Z₂ of thebeam. The transmitted pulse energy will readily reach the second focalzone, since the microbubbles in the nearer first zone were previouslydestroyed. FIG. 14b illustrates this transmission to the second focalzone. Line 282 indicates that the remaining microbubbles at the end ofthe first zone and the beginning of the second will be destroyed by thesecond destruction pulse, as will microbubbles around the focal point asindicated by line 280. Beyond the second focal point F₂ the degree ofmicrobubble destruction will decline as pulse energy declines, asindicated by line 284. A second interrrogation pulse may be transmittedfollowing the second destructive pulse to differentially detect thesecond sequence of microbubble destruction events.

[0056] Similarly, a third destruction pulse is transmitted along thebeam direction, focused at the deepest focal point F₃ in the deepestfocal zone Z₃. The pulse energy readily reaches the third focal zone dueto the earlier depletion of microbubbles at shallower depths. The thirddestruction pulse destroys the remaining microbubbles between the secondand third zones as indicated by line 292 in FIG. 14c, destroysmicrobubbles around the focal point as indicated by line 290, anddestroys a decreasing amount of microbubbles beyond the focal point F₃as indicated by line 294. A third interrogation pulse follows fordifferential detection of the microbubble destruction events in andaround zone Z₃.

[0057] In practice it has been found that peak microbubble destructionis not centered exactly about the focal point axially, but in a depthregion just prior to the focal point. This factor should be taken intoconsideration when considering the placement and overlap of multizonemicrobubble destruction regions.

[0058] The detected destruction events over the three zones are thencombined in accordance with the expression

|P _(F1) −P′ _(F1) |+|P _(F2) −P _(F2) |+|P _(F3) −P′ _(F3)|

[0059] where P_(Fn) represents echoes following a destructive pulsetransmission to a given focal zone and P′_(Fn) represents the echoesfrom a subsequent interrogation pulse. The echoes from each focal zoneare spliced together to form a complete image line to the maximum depthof the image. In a preferred embodiment, instead of just detectingmicrobubble destruction events over the given focal zone, the techniqueconventionally used in multizone focus imaging, echoes are detected overthe full depth following each pulse. This enables the recording ofmicrobubble destruction events outside the given focal zone, providingthe greatest detection of destruction events. Thus, each pulse echo paircontains a line of echoes over the full image depth, which are thencombined to record the maximum number of microbubble destruction eventsfor the full image line.

[0060] It is also seen that, instead of transmitting a pair of pulses tointerrogate each focal zone, the echoes returned from later focal zonetransmissions can be combined with earlier echoes to differentiallydetect destruction events. That is, the first term of the aboveexpression could be |P_(F1)−P_(F2)|, for instance. However, the use ofpulse pairs for each focal zone is preferred, as the aperture changesaccompanying focal zone changes can deleteriously affect the precisionof the technique.

[0061] More uniform, artifact-free multizone microbubble destructionimages can be obtained by pulsing nonadjacent beams with time successivepulses. This ensures that each line of microbubbles will beapproximately uniformly undisturbed at the beginning of the multizonesequence, preventing successions of bright and dim lines in theultrasonic image.

[0062]FIGS. 15a, 15 b and 16 illustrate a preferred technique fordisplaying contrast agent enhanced images when tissue perfusion is beingobserved. FIG. 16 illustrates a cross sectional view of the heart,including the myocardium 260 and the blood pool 250 within a chamber ofthe heart. When a contrast agent has been introduced into thebloodstream, a great quantity of the agent will be contained withinlarge blood pools such as the heart chambers and major vessels, whileonly a relatively small quantity of contrast agent will enter tissue andorgans by way of capillary structures. In the heart image of FIG. 16, alarge quantity of contrast agent will be present in the blood pool 250while a lesser amount will be infused by capillary flow into themyocardium 260.

[0063] A conventional ultrasonic display of the cross sectional image ofFIG. 16 will cause pixels of greater signal level to be illuminated withgreater brightness or color. A typical display mapping characteristicwhich provides this result is shown in FIG. 15a by mappingcharacteristic 252. As detected pixel values increase, the displaypixels are shown with increasing brightness or color until reaching amaximum plateau level. As a result, the blood pool area 250 in FIG. 16will be shown brightly or highly colored, whereas the myocardium 260will be only dimly illuminated or colored.

[0064] When the myocardium is the area of interest in FIG. 16, a displaymapping characteristic such as that shown in FIG. 15b is employed. Thecurve 254 in this drawing is seen to begin at a zero level to suppressnoise in the image, then rises to a high level 256. Thereafter itdeclines to a level 258 for higher detected signal values. As a result,lower detected pixel values will be mapped to brightly illuminated orcolored display pixels, and higher detected pixel values will be mappedto more dimly illuminated or colored display pixel values. As aconsequence of this mapping, the myocardium 260 in FIG. 16 will bebrightly illuminated or colored, while the central blood pool is onlydimly colored or illuminated. This emphasis provides highlighting ofcontrast agent perfused tissue over blood pool areas.

[0065] Pulse transmission techniques can afford further improvement incontrast agent destruction and detection. While the exact physicalmechanisms caused the by interaction of microbubbles with acousticenergy are quite complex, the sizes of microbubbles have an effect upontheir destruction at certain frequencies. Since a microbubble contrastagent is often comprised of microbubbles of a wide range of diameters,microbubble destruction events can be increased by transmitting a chirpor multifrequency pulse. By transmitting a frequency modulated pulse,the probability of transmitting destructive energy for a greater rangeof microbubble sizes is increased. In addition, by modulating both thefrequency and amplitude of the destructive pulse, both microbubbledestruction and controlled oscillation can be induced. The initial highamplitude, low frequency period of the pulse, followed by a loweramplitude, higher frequency period can induce microbubble shelldestruction followed by oscillation of the released microbubble.

[0066] Another transmission technique which affords high pulse rates(PRF) is illustrated in FIGS. 17a-17 c. FIG. 17a illustrates thetransmission of a first pulse P₁ for contrast agent imaging of theheart, followed by a second pulse P₂. In this example the pulses aretransmitted at a low PRF, and a significant period of time existsbetween the transmission times of the pulses. During this time echoes300 are first received from contrast agent in the myocardium, and laterechoes 302 are received from the more distant pericardium.Differentiation of the echoes following the two pulses will detect thepresence of contrast agent in the myocardium, followed by detection ofthe pericardium itself.

[0067] For procedures where it is only desirable to perform contrastagent imaging of the myocardium, a higher PRF transmission can beemployed as shown in FIG. 17b. The higher PRF pulses have theunfortunate result of artifact development. Echoes 300 return from thecontrast agent in the myocardium following pulse P₁. But echoes 302returning from the pericardium in response to the first pulse P₁ appearin the interval following the second pulse P₂ and can manifestthemselves as an artifact in the image when echoes following the twopulses are differentiated. To eliminate the artifact from the laterreturning echoes, incoherent detection is employed prior todifferentiation by the apparatus of FIG. 1. As shown in FIG. 17c,incoherent detection and differentiation results in positive polarityechoes 300′ from the myocardium microbubbles, and negative polarityechoes 302′ from the pericardium. The unwanted negative polarity echoes302′ from the pericardium can then be removed by thresholding orclipping at the baseline, leaving only the desired detection of thecontrast agent in the myocardium.

What is claimed is:
 1. A method for ultrasonically detecting theperfusion rate of tissue by ultrasonic contrast agents comprising:introducing an ultrasonic contrast agent of microbubbles into thebloodstream; transmitting an ultrasonic pulse which significantlydisrupts microbubbles in said tissue; and following the significantdisruption of said microbubbles, ultrasonically measuring the degree ofmicrobubble replenishment of said tissue by acquiring echoes in responseto substantially nondestructive pulses repetitively triggered or gatedin relation to the heart cycle.
 2. The method of claim 1, whereinultrasonically measuring further comprises measuring the degree ofcontrast agent replenishment at a substantially constant point in theheart cycle.
 3. The method of claim 2, wherein ultrasonically measuringfurther comprises acquiring echoes in response to pulses triggered orgated in synchronization with the diastolic phase of the heart cycle. 4.The method of claim 1, further comprising utilizing the acquired echoinformation to produce an ultrasonic image of the perfusion rate oftissue in a region of the body.
 5. The method of claim 4, whereinproducing an ultrasonic image further comprises displaying a portion ofthe region which is fed from a stenotic blood supply with a differentdegree of contrast agent reinfusion than a portion of the region whichis not fed from a stenotic blood supply.
 6. The method of claim 5,wherein displaying a portion of the region which is fed from a stenoticblood supply with a different degree of contrast agent reinfusionfurther comprises displaying a portion of the region which is fed from astenotic blood supply with a different perfusion rate than a portion ofthe region which is not fed from a stenotic blood supply.
 7. A methodfor ultrasonically measuring the perfusion rate of tissue by ultrasoniccontrast agents comprising: introducing an ultrasonic contrast agent ofmicrobubbles into the bloodstream; transmitting a chirp ormultifrequency ultrasonic pulse which significantly disruptsmicrobubbles in said tissue; and following the significant disruption ofsaid microbubbles, ultrasonically measuring the degree of microbubblereplenishment of said tissue.
 8. The method of claim 7, wherein thetransmitted pulse is a frequency modulated pulse.